The dominant x-ray radiographic imaging system formats are the large area (flat) detectors and the slit/slot scanning detectors. Large area x-ray detectors are employed (in most cases) in stationary configurations in a face-on geometry. They tend to be expensive to build due to readout response uniformity problems over large areas. Response uniformity problems also limit the thickness of the detector material (typically a scintillator or semiconductor, although gas detectors have been used for applications such as mammography) that interacts with the incident ionizing radiation. Large area x-ray detectors are often used with an anti-scatter grid which increases patient dose since a fraction of unscattered radiation used for imaging is also removed.
Slit scanning x-ray detectors are line detectors that can be implemented in an edge-on or face-on geometry in a slit scan system. Complexity is introduced by the need for a rapid readout and the need for the uniform motion of the scanning system (translation and rotation motions of the radiation source, detector, and patient are all possible). Response uniformity issues are simplified relative to large area detectors. Slit scan systems typically offer more efficient detectors than large area detectors but they are far less efficient with regard to x-ray tube output utilization. Slit widths on the order of 0.05 mm (50 um) are in use for mammography slit scanning. The first commercial CT scanners used a single slit in which the patient and table were moved through the x-ray beam. Commercial slit scans systems typically require high output x-ray tubes (synchrotron sources are not yet cost-effective). The inefficient utilization of the x-ray tube output can be ameliorated by using multiple slits or a slot (or multiple slots). One limitation of a slot versus a slit is that there is an increase in scatter detection. A slot scan detector has two or more adjacent lines of detectors. Cost is a factor in deciding to implement multiple slit or slot (or multiple slot) configurations that increase the available detector area. If the detector position remains fixed with respect to the slot then a time delay integration (TDI) readout technique is usually employed. A commercial CT scanner with multiple rows of detectors can be viewed as a particular type of slot scan system but the detectors don't operate in a TDI readout mode.
An additional issue to consider is how to treat the detected signal. Integrating the energy of the total detected signal (referred to as energy integration or integration), involves summing the signals from all events over a given period of time. It is the simplest method and widely implemented in large area detector systems and slot scan systems. Depending on the detected beam spectrum the integration approach may represent an undesirable loss of information content (Nelson, U.S. Pat. No. 4,958,368 and U.S. Pat. No. 4,969,175). An alternative approach, photon counting, involves detecting individual x-ray events. A simple photon counting technique typically employs a low signal level threshold cutoff below which events are ignored. A more-sophisticated photon counting technique implements upper and lower signal threshold cutoffs. The most sophisticated photon counting technique involves segregating events into energy bins (providing energy resolution or spectroscopy). In the past the low x-ray energies (typically less than 120 KeV max), the high x-ray event rates encountered in diagnostic medical radiography and CT, as well as a lack of fast scintillators with good conversion efficiency and fast, low-noise photodetectors and readout electronics have been a barrier to utilizing photon counting techniques for scintillator-based slit and slot scan systems. (At least one commercial slit scan system for mammography implements photon counting but it utilizes a silicon semiconductor x-ray detector.) Although photon counting techniques are more expensive to implement than simple integration, cost issues are becoming less of a barrier with the introduction of desirable scintillators, photodetectors, and sophisticated high speed readout electronics within the last few years.
Semiconductor detectors based on semiconductor materials such as Si, Ge, GaAs, CdTe, CdZnTe (CZT), HgI2, PbI2, Se, Diamond, etc. (including films of these materials) have been pursued in radiography since they may offer a low noise, direct electronic readout (the potential for photon counting with energy resolution) as well as acceptable or superior detection efficiency. Si and Ge are widely used in radiation detectors. Both are capable of reasonably fast readout speeds and excellent energy resolution. They can withstand large radiation doses before their properties deteriorate noticeably. Drawbacks may include low atomic number and density (Si) or cooling requirements (Ge is typically cooled with liquid nitrogen). Semiconductor materials may have yield issues for acceptable thicknesses, dead layers, polarization issues, or readout times for electronic signals that may be excessive for slit or slot scanning systems (including CT systems).
Scintillator detectors compete with semiconductor detectors based on cost, readout rates, and desirable material properties. The range of scintillator materials available for nuclear medicine and x-ray imaging has increased significantly since the 1980s in which scintillators such as Gd2O2S:Tb, NaI:Tl, CsI:Na, CsI:Tl, columnar CsI:Na/Tl, CaWO4, CdWO4, BaF2, and BGO played a prominent role. (Gd-based scintillators and LiI:Eu were also useful for neutron detection.) Newer scintillators (including, but not limited to GSO, LSO, LYSO, LuAP, LaBr3, LaCl3, GdI3, LuI3, SrI2, BaHfO3, SrHfO3, PbWO4, and CsI:Tl,Sm) with moderate-to-high density, desirable atomic composition, good light output, fast decays times, and reasonable indices of refraction have been developed for nuclear medicine applications such as probe detectors, gamma cameras, and PET cameras. Ceramic and nano-particle ceramic (and nanocomposite) implementations of nuclear medicine scintillators are being developed and tested. Improved ceramic scintillators are currently being used in medical CT scanners. Efficient manufacturing techniques to build structured 1-D and 2-D scintillator arrays for x-ray (including gamma ray) photon detection have been developed (Nelson, U.S. Pat. No. 5,258,145). Glass, plastic, liquid, and noble gas scintillator materials have also been used for nuclear medicine imaging of gamma rays and particles such as alphas and betas. (Particular versions of these detectors have been used in x-ray radiography as well as for neutron detection and neutron radiography.) High speed photodetector readout in a compact format has become more practical due to development of Si photodetectors such as electron multiplying CCDS (EMCCDs), Geiger-mode silicon photomultiplier (SiPM) arrays, internal discrete amplification detector (iDAD) arrays, avalanche photodiode (APD) arrays or position-sensitive APDs (PSAPDs), etc. that can provide (in some cases) readout times that range from nanoseconds to sub-nanoseconds, and can provide gain. Subgroups of these pixels can share a common output for the total energy of the signal and provide a weighted spatial location if desired (Nelson, Application No. 60/667,824). Silicon drift detectors (SDDs) are capable of low noise readout but require additional amplification. Photoemissive detector formats that offer a high-speed readout have also been developed such as 2-D position sensitive photomultiplier tubes or PSPMTs, 2-D hybrid photoemissive-photodiode arrays (Braem A., et al., Nuc. Instr. Meth. Phys. Res. A Vol. 525, pp. 268-274, 2004 and Vol. 580, pp. 1513-1521, 2007), scintillator-based intensifiers optically coupled to photodetector arrays, microchannel plate amplifiers (microchannel plates) coupled to photodetector or metal arrays, etc. In addition, the conversion efficiencies of several photocathode materials have improved such that they are competitive with silicon photodetectors for the shorter wavelengths encountered with many fast-decay scintillators.